This invention relates to magnetic resonance (MR) imaging.
A prior art magnetic resonance imaging apparatus is shown in FIG. 1. A patient 1 (shown in section) is slid axially into the bore 2 of a superconducting magnet 3, and the main magnetic field is set up along the axis of the bore, termed by convention the Z-direction. Magnetic field gradients are set up, for example, in the Z-direction, to confine the excitation of magnetic resonant (MR) active nuclei (typically hydrogen protons in water and fat tissue) to a particular slice in the Z-direction e.g. that illustrated in FIG. 1 and, in the horizontal X and the vertical Y-directions as seen in FIG. 1, to encode the resonant MR nuclei in the plane of the slice. An r.f. transmit coil (not shown) applies an excitation pulse to excite the protons to resonance, and an r.f. receive coil array consisting of a pair of coils 4, 5 picks up relaxation signals emitted by the disturbed protons.
To encode/decode received signals in the Y-direction, the signals are detected in the presence of a magnetic field gradient, termed a frequency encode or read-out (R.O.) gradient, to enable different positions of relaxing nuclei to correspond to different precession frequencies of those nuclei about the direction of the main magnetic field due to the influence of the gradient. The data is digitised, and so for each r.f. excitation pulse, a series of digital data points are collected, and these are mapped into a spatial frequency domain known as k-space (FIG. 2). Each r.f. pulse permits at least one column of digital data points to be collected.
To encode/decode the received signals in the X-direction, after each r.f. pulse has been transmitted and before data is collected with the read-out gradient applied, a magnetic field gradient in the X-direction is turned on and off. This is done for a series of magnitudes of magnetic field gradients in the X-direction, one r.f. pulse typically corresponding to a different magnitude of gradient in the X-direction. The series of measurements enable spatial frequencies to be built up in the X-direction.
On the k-space matrix shown in FIG. 2, the columns of data points correspond to data collected at different magnitudes of phase-encode (P.E.) gradients.
The field of view imaged by the magnetic resonance imaging apparatus depends on the spacing of the data points in the phase-encode and read-out directions, and the resolution of the image depends on how far the points extend in each direction i.e. how large the maximum phase-encode gradient is, and on the magnitude of the read-out gradient combined with the duration of data collection.
Conventionally, the data collected by the r.f. receive coil arrangement and depicted in FIG. 2 is subject to a two dimensional fast Fourier Transform in a Fourier Transform processor (not shown) to produce a pixelated spatial image.
A slice image is shown in FIG. 3. For the purposes of explanation, the symbol of a circle 1a, has been illustrated in both the patient 1 shown in FIG. 1 and the image shown in FIG. 3. FIG. 3 implies that the spacing of data points in the phase-encode gradient direction is sufficient to image the whole of the circle shown in FIG. 1.
Between each r.f pulse, there is a certain minimum pulse repetition time, and the collection of data implied by FIGS. 2 and 3 may therefore take an undesirably long time.
One technique used to reduce the data collection time is to cut out, say, half the phase-encode steps e.g. by keeping the same maximum phase-encode gradient but omitting every other column of data. This would then halve the data collection time.
The spacing of the data points in the phase-encode direction would now have doubled, so that the field of view in the corresponding image domain would have halved. (The field of view in the read-out direction would remain the same because the number of data points collected during read-out would remain the same.) The imaged area would now cover little more than half the width of the circle illustrated in FIG. 1. This is shown by the area 1b in FIG. 5. Unfortunately, aliasing causes the regions at the side of the circle to be folded back into the half-width area, the left hand region in FIG. 5 corresponding to the right hand region of the image, and vice versa.
To enable the data to be unfolded, the data is acquired using parallel imaging.
Parallel imaging makes use of spatial sensitivity differences between individual coils in an array to reduce the gradient encoding required during image acquisition. This reduces acquisition times by decreasing the number of phase-encoded lines of k-space that must be acquired. There are three distinct classes of practical implementation of parallel imaging, which are known as SENSE (Magnetic Resonance in Medicine 42: 952-962 (1999)—SENSE: Sensitivity Encoding for Fast MRI by Klaas P Pruessmann, Markus Weiger, Markus B Scheidegger and Peter Boesiger), SMASH (WO-A-98/21600 and Magnetic Resonance in Medicine 38: 591-603 (1997)—Simultaneous Acquisition of Spatial Harmonics (SMASH): Fast Imaging with Radiofrequency Coil Arrays by Daniel K Sodickson and Warren J Manning) and SPACE-RIP (WO-A-00/72050 and Magnetic Resonance in Medicine 44:301-308 (2000)—Sensitivity Profiles from an Array of Coils for Encoding and Reconstruction in Parallel (SPACE RIP) by Walid E Kyriakos, Laurence P Panyah, Daniel F Kaches, Carl-Frederick Westin, Sumi M Bao, Robert V Mulkern and Ferenc A Jolesz). All of these methods require information about the coil sensitivity profiles (reference data), which is used to regenerate a full image data set from the sub-sampled k-space acquisition (target data).
SENSE operates in the image domain for both the target image data and the coil reference data. The method can be used with a wide range of coil geometries. A typical receive coil arrangement comprises coils 4 and 5 placed on opposite sides of the patient arranged in FIG. 1, in order that they have different fields of view. The target data is acquired for each receive coil with a reduced field of view, which results in aliasing, so that each coil produces a k-space representation as shown in FIG. 4, which can be Fourier Transformed into an aliased image as shown in FIG. 5. The two aliased images of FIG. 5 are then unfolded to the full field of view on a pixel by pixel basis using reference data, which records the relative responses of the receive coils 4 and 5. Reduced field of view imaging imposes a requirement of uniformly spaced samples in the phase-encode direction in k-space. Since processing concerned with unfolding is done in the image domain, individual pixels in the reduced field of view data get unfolded by integer numbers of final pixels (i.e. 1→1, 1→2, 1→3 etc). This requires solution of a set of linear simultaneous equations in which pixel intensities are weighted by the coil sensitivity at the final pixel locations. The numerical condition of these equations determines the local noise properties of the unfolded image, so that the signal-to-noise ratio (SNR) varies from pixel to pixel. The signal-to-noise ratio is better in the regions (e.g. in FIG. 5) where no aliasing occurs than where it does occur. The resulting patterns of noise variation generally reflect the coil geometry and can have a strong perceptual effect.
SPACE RIP uses k-space target data as input in conjunction with a real space representation of the coil sensitivities to directly compute a final image domain output, that is, the Fourier transform is embedded into the matrix involved. An unfolded image is directly produced from the reduced phase-encode gradient encoded collected data for the coils of the array (FIG. 4). Thus, it is a hybrid k-space/real space method and has a higher computational burden than either SENSE or SMASH, but does not require uniform sampling of k-space.
SMASH operates in k-space for the target image data but uses a real space representation of the coil sensitivity profiles. SMASH employs linear combinations of the coil reference data to explicitly construct spatial harmonics that are required to synthesis missing k-space lines. It does not suffer from spatially varying signal-to-noise ratio in the final images, since each point in k-space contributes to the whole image in the image domain.
A typical coil arrangement for SMASH is shown in FIG. 6. An array of coils 6 to 13 is arranged beneath the spine 14 of a patient 7 (shown schematically). Such a coil arrangement can be used to produce a saggital (vertical longitudinal) section through the spine (plane 15). The response patterns of the individual coils is shown in FIG. 7. If the outputs of the individual coils is suitably weighted and summed, it can be seen that, for example, the response of FIG. 8, and its simplified form of FIG. 9, can be produced. Such a weighted and summed signal modulates received r.f. signals along the length of the array in the same way as a phase-encoding gradient in the Z-direction modulates r.f. signals received by an equivalent received coil 16 (shown dotted). Accordingly, SMASH uses weighted combinations of the outputs of the individual coils of the array to simulate the effect of phase-encode gradients on the received r.f signals. The fundamental is shown in FIGS. 8 and 9, but different weightings can be used to produce higher harmonics. Thus, signals representing several phase-encode gradient lines can be produced for the application of one phase-encode gradient.
However, SMASH is somewhat restrictive in the coil geometries it can accommodate. In particular, it is not well suited to use with very few coils and the requirement to generate specific spatial harmonics necessitates a given relationship between the imaging field of view and the coil structure.
A simplified example of the SMASH process will be described with reference to FIGS. 10 and 11. The vertical lines represent columns of points in k-space in the frequency encode direction.
It is assumed that the spine array of FIG. 6 is used, but that there are four coils. It is assumed that there are only four phase-encoding gradient steps. Thus, each coil collects signal at each phase-encode gradient so that 16 signals are collected. The outputs of each coil for each phase-encode gradient are combined to produce a resultant signal for that phase-encode gradient. For example, in FIG. 10, line 17 represents the output of each coil at the second phase-encode gradient, and these signals are used to produce a resultant signal at that phase-encode gradient. The interpolated signals corresponding to notional phase-encode gradients are produced from these resultant signals together with a phase rotation reflecting the different modulation between the phase-encode gradient of the resultant and of the interpolated phase-encode line. The resultant and the interpolated lines in FIGS. 10 and 11 are produced on a point by point basis from the resultants of the corresponding points on the lines 17.
Clearly there will usually be considerably more than four phase-encode gradients, and more coils than four may be used. Nevertheless, it will be seen that in the SMASH method, measurements may be made as a reduced number of phase-encode gradients, and data corresponding to the omitted ones may be generated using the SMASH method.